Dental x-rays are typically taken with a film that is placed in the patient's mouth. The film is exposed through the teeth by an x-ray source that resides outside the patient's head. While this method has been in use for many years, it has a number of disadvantages. First, the patient is exposed to a significant dose of x-rays. This dose is accumulative over the patient's lifetime. Second, the time, cost, and equipment needed to process the film increases the cost of the dental examination. Third, the chemicals utilized in processing the film pose a disposal problem.
These problems have led to several attempts to replace the film component of the traditional x-ray examination with a solid-state sensor that is placed in the patient's mouth to record the x-ray image. In such systems, a layer of scintillation material is used to convert the x-rays to visible light. The visible light is then imaged onto a solid-state imaging array. Since solid-state x-ray sensors of this type are significantly more sensitive to x-rays than the films utilized today, the x-ray dosage can be reduced by typically a factor of 10. In addition, the sensor is re-used, and hence, the environmental problems associated with the conventional x-ray system are avoided. Finally, since the image is in digital form, systems based on solid-state sensors are easily adapted to paperless office systems.
CMOS imaging arrays typically include an array of pixel elements in which each element has a photodiode and an active gain stage. The photodiode generates and stores a charge that is related to the amount of light that was received by the photodiode during a predetermined exposure period. The active gain stage typically converts this charge to a voltage that is readout on a bus to a readout circuit that digitizes the voltage to provide the intensity value associated with the pixel in question.
In a conventional photographic application, the physical size of each pixel is set by considerations that are more or less independent of the images that are to be captured, since the camera using the imaging array includes a lens that matches the image to the size of the array. Hence, the conversion of a conventional film-based camera to a digital camera can be made with relatively few changes to the camera. For example, if the imaging array has a size that differs from that of the film that is being replaced, the magnification of the lens system can be changed to assure that the image covers the imaging array. In such applications, the number of pixels and the sensitivity of the array are the parameters of interest. Arrays with larger numbers of pixels provide images with finer detail. Similarly, arrays with higher sensitivity can be utilized in lower light situations. The cost of the array is determined by the size of the die on which the imaging array is constructed, larger dies being more expensive. The sensitivity of the array depends on the amount of silicon in each pixel that is devoted to the photodiode, as opposed to the active gain stage and other circuitry. Hence, cost, resolution, and sensitivity are traded against one another to arrive at an acceptable design.
In contrast, in dental applications, the size of the die is fixed by the geometry of the patient's mouth. A dental x-ray image is essentially the shadow of the teeth on the imaging surface. Hence, the imaging array must be large enough to capture the same area as the conventional x-ray film without any additional lens to compensate for size differences. The required resolution is likewise set by the x-ray imaging process, which has an inherent blurring function built into it. Hence, once the pixel size is below some threshold size that depends on the blurring, no significant improvement in image quality is obtained by further reducing the size of the pixels in the array. As a result of these considerations, the optimum pixel would be a square with a side of approximately 25 microns.
The preferred light-sensing element in CMOS imaging arrays is a “pinned” photodiode. The diode is doped such that the charge storage region of the photodiode is at a potential that is significantly higher than the input to the active circuitry that converts the stored charge to a voltage. This arrangement assures that all of the charge accumulated during the period in which the photodiode is exposed to light is removed during readout and reset processes. If any charge were to remain, the next image taken by the array could include a ghost of the previous image.
Unfortunately, constructing the pinned photodiodes of the desired size for x-ray imaging in conventional CMOS processes is difficult. Hence, conventional CMOS imaging arrays having larger numbers of pixels of a smaller size are used. In effect, the 25×25 micron area is broken up into a number of smaller pixels of a size that can be constructed in CMOS. The results from these pixels are then added together after the image is formed to provide an image that approximates the image that would have been formed using the larger pixel size.
Unfortunately, this approach has a number of problems. First, each pixel includes an active gain element and the gate circuitry associated with reading out the individual pixels on the readout buses. The added circuitry reduces the fill-factor of the pixel, i.e., the ratio of the photodiode area to the pixel area. Hence, sensitivity is lost, which leads to increased x-ray exposure times. Second, the readout time is increased. If the 25 micron pixels are broken into 5 micron pixels, then there are 25 times more pixels that must be readout. If the array is organized as a rectangular array with rows and columns of pixels, the number of rows is increased by a factor of 5, and hence, the readout time is increased by a factor of 5 even if an entire row is read in parallel by providing an analog-to-digital converter for each column of pixels. In addition, the increase in the number of columns leads to a significant increase in the number of ADCs needed to digitize the image, which further increases the cost of the dental sensor.